1. Technical Field of the Invention
The present invention relates to a magnetic resonance imaging apparatus and magnetic resonance imaging data processing method that can reduce positional shift of tomographic image, poor fat suppression, etc. caused by a shift in the spatial distribution of intensity (preferably uniform) of static magnetic field Bo as it varies with temperature variation of a thermally coupled gradient magnetic field coil unit.
2. Description of the Related Art
Conventionally, it is a practice to use a magnetic resonance imaging (MRI) apparatus 1 as shown in FIG. 15 (see JP-A-2002-85369, for example), as a monitoring apparatus at the medical site.
An MRI apparatus 1 has a static-magnetic-field magnet 2 cylindrical in form for forming a static magnetic field within which can be laid an examination subject P having an imaging region where a gradient magnetic field is formed in directions of X, Y and Z axes by means of gradient magnetic field coils 3x, 3y, 3z of a gradient magnetic field coil unit 3 so that magnetic resonance can be caused with the nuclear spin within the examination subject P by transmitting an RF (radio frequency) signal from an RF coil 4. By utilizing a nuclear magnetic resonant signal (NMR) caused by excitation, an image can be reconstructed as to the examination subject P.
Namely, a static magnetic field is formed within the static-magnetic-field magnet 2 by a static-magnetic-field power supply 5. Furthermore, according to an instruction from an input device 6, a sequence-controller control unit 7 provides a sequence, as signal control information, to a sequence controller 8. The sequence controller 8 in turns controls a gradient magnetic field power supply 9 connected to the gradient magnetic field coils 3x, 3y, 3z and a transmitter 10 for providing an RF signal to the RF coil 4, according to a sequence. Due to this, a gradient magnetic field is formed at the imaging region so that an RF signal can be transmitted to the examination subject P.
In this case, the X-axis gradient magnetic field, the Y-axis gradient magnetic field and the Z-axis gradient magnetic field, formed by the gradient magnetic field coils 3x, 3y, 3z, are used mainly as a magnetic field for phase encoding (PE), a magnetic field for readout (RO) and a magnetic field for slice selection (SL), respectively. For this reason, the X, Y and z coordinates, defining nuclear positional information, are respectively converted into a phase, a frequency and a slice region of nuclear spin.
The NMR signal, caused by an excitation of nuclear spin at the interior of the examination subject P, is received at the RF coil 4 and provided to a receiver 11 where it is digitized and converted into raw data. Furthermore, the raw data is taken in the sequence-controller control unit 7 through the sequence control 8. The sequence-controller control unit 7 arranges the raw data in k-space (Fourier space) formed in a raw-data database 12. Then, an image reconstruction unit 13 performs a Fourier transform on the raw data arranged in the k-space and provides it to a display device 14 to make a display, thereby reconstructing an image of examination subject P.
For such MRI apparatuses 1, there is a recent attempt to increase the imaging rate by forming a intense gradient magnetic field at increased rate. The representative approaches of high-speed imaging based on the MRI apparatuses 1 include a known scheme of EPI (echo planer imaging) capable of acquiring a plurality of echoes by one excitation. In the EPI scheme of imaging, the gradient magnetic field power supply 9 and the transmitter 10 are controlled by the sequence controller 8 according to an EPI sequence set up based on the EPI scheme.
In the case of imaging according to an EPI sequence, a PE gradient magnetic field is formed with high intensity by providing high power electrical current to the gradient-magnetic-field coil unit 3 in order to improve frequency resolution in the direction of PE gradient magnetic field.
However, in the case of imaging according to the usual EPI scheme, there is a problem of the influence of magnetic field non-uniformity with the result of deteriorated image quality. To avoid this, a current is supplied from a shim-coil power supply 16 to a cylindrical shim coil 15 provided coaxially to and inward of a magnet forming the static magnetic field, thereby making the static magnetic field uniform.
Conventionally, a ferrous shim is inserted in the gradient magnetic field coil 3x, 3y, 3z of the MRI apparatus 1 in order to correct for non-uniformity of the static magnetic field. However, when the sequence is continuously executed in order to carry out dynamic image scanning at successive points in time of the examination subject P, heat is generated on the gradient-magnetic-field coil 3x, 3y, 3z to thereby raise its temperature with the passage of image-taking time. Consequently, the heat from the gradient-magnetic-field coil 3x, 3y, 3z conducts to the ferrous shim, thus raising the temperature of the ferrous shim and lowering the magnetic permeability of the gradient-magnetic-field coil 3x, 3y, 3z. This causes so-called a Bo shift in that the magnetic flux density passing through the ferrous shim decreases to make spatially non-uniform the static magnetic field, i.e. magnetic flux density at the imaging region. The Bo shift, while predominant in the zero order component, also causes higher-order components.
Furthermore, according to the Larmor formula, the resonant frequency of the MNR signal is proportional to magnetic field intensity. Thus, resonant frequency is varied by a Bo shift, raising such problems as shifting the apparent position of a tomographic image obtained by reconstruction and deteriorating image quality due to a deformed tomographic image. Particularly, in a forced load test, there arises a case that resonant frequency can vary by 100 Hz or more.
Meanwhile, there are conventional cases pointing out the problem that it is not sufficient to suppress (fat suppression) of an NMR signal caused by fat molecules.
Such a positional shift of tomographic image and poor fat suppression caused by a Bo shift is to appear conspicuously in such a case as performing a dynamic scanning consuming power over a long time (10 minutes or longer) as encountered particularly in an EPI sequence. Furthermore, there is a recent tendency for the problem of Bo shift to become more conspicuous because of the execution of a long-time imaging of an FE-EPI scheme using an EPI scheme together with an FE (field echo) scheme that causes an NMR signal by inverting the gradient magnetic field.
Recently, there is a dissemination of the multi-channel phased-array coil structured with a plurality of coils. Where the phased-array coil is employed as an RF coil 4, the RF coil 4 has a sensitivity varying from channel to channel in accordance with the spatial distribution of a Bo shift. This raises a problem of causing a variation between the correction values for correcting the sensitivities on the channels of the RF coil 4 because of correspondence with the sensitivity distribution on the RF coil 4.
Meanwhile, the gradient-magnetic-field coil unit 3 constitutes a major source of noise generation in the MRI apparatus 1. Utilized is an MRI apparatus 1 having a silencer mechanism utilizing a vacuum technique in the gradient-magnetic-field coil 3x, 3y, 3z of the gradient magnetic field coil unit 3. When the vacuum degree in the silencer mechanism changes, a phenomenon occurs that is similar to the case of temperature rise on the ferrous shim.
For this reason, considerations have been made on possible corrective measures including a method to remove the ferrous shim from the magnetic field coil 3x, 3y, 3z, a method to provide a cooling function so as not to raise the temperature of the ferrous shim, and so on. However, such approaches lead to cost increases.
Besides, there are differences in the effect of temperature rise of ferrous shim upon an image between MRI apparatuses 1 in accordance with the use amount and layout of a ferrous shim. This makes countermeasures further difficult.